Visible images of the results of exposing a patient or object of interest (hereafter "subject") to x-radiation or other high energy radiation (referred to collectively herein as x-radiation or x-rays or the like) can be divided into two broad categories: immediate images and delayed images. Immediate images are commonly producing by impinging transmitted or reflected x-radiation at a screen containing a prompt emission radiographic phosphor. The radiographic phosphor emits light in response to the x-radiation. This procedure, in medical x-ray imaging, is commonly referred to as "fluoroscopy". Delayed images can be produced by directing the x-radiation at photographic film. The shortcoming of this approach is that the required x-radiation dosage is quite high. An alternative approach is to place a sheet of photographic film against a screen containing prompt emission radiographic phosphor and to direct the x-radiation at the screen. The screen is costly, but has a relatively long useful life. The photographic film characteristics are tailored to complement the emission produced by the radiographic phosphor. One shortcoming of this approach is that the film and screen must be enclosed within a cassette or other light-tight enclosure to prevent ambient light from exposing the film. Another shortcoming is the broadening of image features that occurs as a result of the spatial separation of a first plane at which the x-ray image is absorbed and the light image generated and a second plane at which the light image is absorbed. The spatial separation can also be exacerbated by localized gaps between the screen and film due to dust particles or irregularities in the cassette. Another process used for producing delayed x-radiation images is known as computed radiography. A x-ray image is captured on a screen containing a storage radiographic phosphor. The latent image on the storage screen is then exposed to a rastorized beam of radiation which causes the storage phosphor to emit radiation on another wavelength. The emitted radiation is captured and displayed or used to print an image. A particular shortcoming of computed radiography is the complexity of the radiation capture system.
As a result of the above shortfalls, many efforts have been made to replace conventional systems with detectors that can provide an immediate digital image. One type of detector uses a technique called, "line scanned radiography". In this technique, x-rays are collimated through a small slit to produce an x-ray fan beam, which after passing through the subject impinge on a detector strip. Scintillators or x-ray phosphors (hereafter collectively referred to as "phosphors") are positioned in the detector strip between the x-ray beam and a photodiode or photosensor. "Design and Characteristics of a Digital Chest Unit", Digital Radiography, SPIE Vol. 314, (1981), pp. 160-165, teaches such a detector strip having a one-dimensional linear silicon photodiode array. While this approach is cost effective, it has the shortcoming that the fan beam and detector, or alternatively the subject, must be translated to produce a two-dimensional image.
U.S. Pat. No. 5,254,480 to Nang T. Tran discloses a process for producing a solid state radiation detector. The detectors produced have an intensifying screen or layer of prompt emitting phosphor overlying a photosensitive layer. In one embodiment, the photosensitive layer is an array of photodiodes. In another embodiment, the photosensitive layer has a pixellated electrode overlying a continuous photoconductor layer and a second electrode. These detectors suffer from the shortcoming that the light emitted by the phosphor is subject to lateral spread after it is emitted from the phosphor layer, resulting in image blur or reduced resolution. The photosensor based detectors just described also tend to suffer additional problems including: very high cost of manufacture; extreme physical fragility; low resistance to damage from x-ray radiation; and configurational complexity.
A variety of two-dimensional digital x-ray detectors are known. An x-ray detector employing an amorphous selenium detector is described in "X-ray imaging using amorphous selenium: "A photoinduced discharge method for digital mammography", J. A. Rowlands et al., Med. Phys., Vol. 18, (1991) p. 421. In this system, amorphous selenium plates are employed as the x-ray detector and the electrostatic image formed on the plate is then read out electronically. This detector is complex, and utilizes selenium which is relatively inefficient in capturing x-rays.
U.S. Pat. No. 5,313,066 to Lee et al discloses an x-ray detector having a photoconductive layer that is responsive to both actinic and x-ray radiation. The photoconductive layer includes a particulate photoconductor dispersed in a photoconductive polymer. The particulate photoconductor is employed as an x-ray absorber which converts the x-rays into electron-hole pairs which then transport through the photoconductive polymer for detection. (The photoconductive polymer itself is minimally absorptive of x-radiation.) This detector has the shortcomings that photoconductors tend to be inefficient x-ray absorbers and transfer of electron-hole pairs from the particulate photoconductor to the polymeric photoconductor is subject to a variety of difficult constraints including the work function necessary for the transfer and the negative effects of any inhomogeneities at the interface between the particulate and polymeric photoconductors.
It would thus be desirable to provide an improved x-ray detector, detector assembly and method that can exhibit both good x-ray absorption efficiency and minimizes lateral spread of light emitted by the phosphor.